Breathing gas delivery method and apparatus

ABSTRACT

An improved methodology and systems for delivery of breathing gas such as for the treatment of obstructive sleep apnea through application of alternating high and low level positive airway pressure within the airway of the patient with the high and low airway pressure being coordinated with the spontaneous respiration of the patient, and improved methods and apparatus for triggering and for leak management in such systems.

CROSS-REFERENCE TO RELATED APPLICATIONS

This is a Continuation of U.S. patent appln. Ser. No. 08/823,855 filedMar. 25, 1997, now U.S. Pat. No. 6,029,664; which is a Continuation ofU.S. patent appln. Ser. No. 08/349,634 filed Dec. 2, 1994, now U.S. Pat.No. 5,632,269; which is a Continuation-in-part of U.S. appln. Ser. No.07/947,156 filed Sep. 18, 1992, now U.S. Pat. No. 5,433,193; which is aContinuation-in-part of U.S. patent appln. Ser. No. 07/411,012 filedSep. 22, 1989, now U.S. Pat. No. 5,148,802.

BACKGROUND OF THE INVENTION

The sleep apnea syndrome, and in particular obstructive sleep apnea,afflicts an estimated 4% to 9% of the general population and is due toepisodic upper airway obstruction during sleep. Those afflicted withobstructive sleep apnea experience sleep fragmentation and intermittent,complete or nearly complete cessation of ventilation during sleep withpotentially severe degrees of oxyhemoglobin unsaturation. These featuresmay be translated clinically into debilitating daytime sleepiness,cardiac disrhythmias, pulmonary-artery hypertension, congestive heartfailure and cognitive dysfunction. Other sequelae of sleep apnea includeright ventricular dysfunction with cor pulmonale, carbon dioxideretention during wakefulness as well as during sleep, and continuousreduced arterial oxygen tension. Hypersomnolent sleep apnea patients maybe at risk for excessive mortality from these factors as well as from anelevated risk for accidents such as while driving or operating otherpotentially dangerous equipment.

Although details of the pathogenesis of upper airway obstruction insleep apnea patients have not been fully defined, it is generallyaccepted that the mechanism includes either anatomic or functionalabnormalities of the upper airway which result in increased air flowresistance. Such abnormalities may include narrowing of the upper airwaydue to suction forces evolved during inspiration, the effect of gravitypulling the tongue back to appose the pharyngeal wall, and/orinsufficient muscle tone in the upper airway dilator muscles. It hasalso been hypothesized that a mechanism responsible for the knownassociation between obesity and sleep apnea is excessive soft tissue inthe anterior and lateral neck which applies sufficient pressure oninternal structures to narrow the airway.

The treatment of sleep apnea has included such surgical interventions asuvalopalatopharyngoplasty, gastric surgery for obesity, andmaxillo-facial reconstruction. Another mode of surgical interventionused in the treatment of sleep apnea is tracheostomy. These treatmentsconstitute major undertakings with considerable risk of post-operativemorbidity if not mortality. Pharmacologic therapy has in general beendisappointing, especially in patients with more than mild sleep apnea.In addition, side effects from the pharmacologic agents that have beenused are frequent. Thus, medical practitioners continue to seeknon-invasive modes of treatment for sleep apnea with high success ratesand high patient compliance including, for example in cases relating toobesity, weight loss through a regimen of exercise and regulated diet.

Recent work in the treatment of sleep apnea has included the use ofcontinuous positive airway pressure (CPAP) to maintain the airway of thepatient in a continuously open state during sleep. For example, U.S.Pat. No. 4,655,213 and Australian patent AU-B-83901/82 both disclosesleep apnea treatments based on continuous positive airway pressureapplied within the airway of the patient.

Also of interest is U.S. Pat. No. 4,773,411 which discloses a method andapparatus for ventilatory treatment characterized as airway pressurerelease ventilation and which provides a substantially constant elevatedairway pressure with periodic short term reductions of the elevatedairway pressure to a pressure magnitude no less than ambient atmosphericpressure.

Publications pertaining to the application of CPAP in treatment of sleepapnea include the following:

1. Lindsay, DA, Issa FG, and Sullivan C.E. “Mechanisms of SleepDesaturation in Chronic Airflow Limitation Studied with Nasal Continuouspositive Airway pressure (CPAP)”, Am Rev Respir Dis, 1982; 125: p.112.

2. Sanders M H, Moore S E, Eveslage J. “CPAP via nasal mask: A treatmentfor occlusive sleep apnea”, Chest, 1983; 83: pp. 144-145.

3. Sullivan C E, Berthon-Jones M, Issa F G. “Remission of severeobesity-hypoventilation syndrome after short-term treatment during sleepwith continuous positive airway pressure”, Am Rev Respir Dis, 1983; 128:pp. 177-181.

4. Sullivan C E, Issa F G, Berthon-Jones M, Eveslage. “Reversal ofobstructive sleep apnea by continuous positive airway pressure appliedthrough the nares”, Lancet, 1981; 1: pp. 862-865.

5. Sullivan C E, Berthon-Jones M. Issa F G. “Treatment of obstructiveapnea with continuous positive airway pressure applied through thenose”, Am Rev Respir Dis, 1982; 125: p.107. Annual Meeting Abstracts.

6. Rapoport D M, Sorkin B, Garay S M, Goldring R H. “Reversal of the‘pickwickian Syndrome’ by long-term use of nocturnal nasal-airwaypressure”, N Engl J. Med, 1982; 307: pp.931-9:33.

7. Sanders M H, Holzer B C, pennock B E. “The effect of nasal CPKP onvarious sleep apnea patterns”, Chest, 1983; 84: p.336.

Presented at the Annual Meeting of the American College of Chestphysicians, Chicago Ill., October 1983.

Although CPAP has been found to be very effective and well accepted, itsuffers from some of the same limitations, although to a lesser degree,as do the surgical options; specifically, a significant proportion ofsleep apnea patients do not tolerate CPAP well. Thus, development ofother viable non-invasive therapies has been a continuing objective inthe art.

BRIEF SUMMARY OF THE INVENTION

The present invention contemplates a novel and improved method fortreatment of sleep apnea as well as novel methodology and apparatus forcarrying out such improved treatment method. The invention contemplatesthe treatment of sleep apnea through application of pressure at variancewith ambient atmospheric pressure within the upper airway of the patientin a manner to promote dilation of the airway to thereby relieve upperairway occlusion during sleep.

In one embodiment of the invention, positive pressure is appliedalternately at relatively higher-and lower pressure levels within theairway of the patient so that the pressure-induced force applied todilate the patient's airway is alternately a larger and a smallermagnitude dilating force. The higher and lower magnitude positivepressures are initiated by spontaneous patient respiration with thehigher magnitude pressure being applied during inspiration and the lowermagnitude pressure being applied during expiration.

The invention further contemplates a novel and improved apparatus whichis operable in accordance with a novel and improved method to providesleep apnea treatment. More specifically, a flow generator and anadjustable pressure controller supply air flow at a predetermined,adjustable pressure to the airway of a patient through a flowtransducer. The flow transducer generates an output signal which is thenconditioned to provide a signal proportional to the instantaneous flowrate of air to the patient. The instantaneous flow rate signal is fed toa low pass filter which passes only a signal indicative of the averageflow rate over time. The average flow rate signal typically would beexpected to be a value representing a positive flow as the system islikely to have at least minimal leakage from the patient circuit (e.g.small leaks about the perimeter of the respiration mask worn by thepatient). The average flow signal is indicative of leakage because thesummation of all other components of flow over time must be essentiallyzero since inspiration flow must equal expiration flow volume over time,that is, over a period of time the volume of air breathed in equals thevolume of the gases breathed out.

Both the instantaneous flow signal and the average flow rate signal arefed to an inspiration/expiration decision module which is, in itssimplest form, a comparator that continually compares the input signalsand provides a corresponding drive signal to the pressure controller. Ingeneral, when the instantaneous flow exceeds average flow, the patientis inhaling and the drive signal supplied to the pressure controllersets the pressure controller to deliver air, at a preselected elevatedpressure, to the airway of the patient. Similarly, when theinstantaneous flow rate is less than the average flow rate, the patientis exhaling and the decision circuitry thus provides a drive signal toset the pressure controller to provide a relatively lower magnitude ofpressure in the airway of the patient. The patient's airway thus ismaintained open by alternating higher and lower magnitudes of pressurewhich are applied during spontaneous inhalation and exhalation,respectively,

As has been noted, some sleep apnea patients do not tolerate standardCPAP therapy. Specifically, approximately 25% of patients cannottolerate CPAP due to the attendant discomfort. CPAP mandates equalpressures during both inhalation and exhalation. The elevated pressureduring both phases of breathing may create difficulty in exhaling andthe sensation of an inflated chest. However, we have determined thatalthough both inspiratory and expiratory air flow resistances in theairway are elevated during sleep preceding the onset of apnea, theairway flow resistance may be less during expiration than duringinspiration. Thus it follows that the bi-level CPAP therapy of ourinvention as characterized above may be sufficient to maintainpharyngeal patency during expiration even though the pressure appliedduring expiration is not as high as that needed to maintain pharyngealpatency during inspiration. In addition, some patients may haveincreased upper airway resistance primarily during inspiration withresulting adverse physiologic consequences. Thus, our invention alsocontemplates applying elevated pressure only during inhalation thuseliminating the need for global (inhalation and exhalation) increases inairway pressure. The relatively lower pressure applied during expirationmay in some cases approach or equal ambient pressure. The lower pressureapplied in the airway during expiration enhances patient tolerance byalleviating some of the uncomfortable sensations normally associatedwith CPAP.

Under prior CPAP therapy, pressures as high as 20 cm H2O have beenrequired, and some patients on nasal CPAP thus have been needlesslyexposed to unnecessarily high expiratory pressures with the attendantdiscomfort and elevated mean airway pressure, and theoretic risk ofbarotrauma. Our invention permits independent application of a higherinspiratory airway pressure in conjunction with a lower expiratoryairway pressure in order to provide a therapy which is better toleratedby the 25% of the patient population which does not tolerate CPAPtherapy, and which may be safer in the other 75% of the patientpopulation.

As has been noted hereinabove, the switch between higher and lowerpressure magnitudes can be controlled by spontaneous patientrespiration, and the patient thus is able to independently governrespiration rate and volume. As has been also noted, the inventioncontemplates automatic compensation for system leakage whereby nasalmask fit and air flow system integrity are of less consequence than inthe prior art. In addition to the benefit of automatic leakcompensation, other important benefits of the invention include lowermean airway pressures for the patient and enhanced safety, comfort andtolerance.

It is accordingly one object of the present invention to provide a noveland improved method for treatment of sleep apnea.

A further object of the invention to provide a novel and improvedapparatus which is operable according to a novel methodology in carryingout such a treatment for sleep apnea.

Another object of the invention is to provide a method and apparatus fortreating sleep apnea by application of alternately high and lowmagnitudes of pressure in the airway of the patient with the high andlow pressure magnitudes being initiated by spontaneous patientrespiration.

Another object of the invention is to provide an apparatus forgenerating alternately high and low pressure gas flow to a consumer ofthe gas with the higher and lower pressure flows being controlled bycomparison of the instantaneous flow rate to the gas consumer with theaverage flow rate to the consumer, which average flow rate may includeleakage from the system, and whereby the apparatus automaticallycompensates for system leakage.

These and other objects and further advantages of the invention will bemore readily appreciated upon consideration of the following detaileddescription and accompanying drawings, in which:

FIG. 1 is a functional block diagram of an apparatus according to theinstant invention which is operable according to the method of theinstant invention;

FIG. 2 is a functional block diagram showing an alternative embodimentof the invention;

FIG. 3 is a functional block diagram of the Estimated Leak Computer ofFIG. 2;

FIG. 4 is a frontal elevation of a control panel for the apparatus ofthis invention;

FIG. 5 is a trace of patient flow rate versus time pertaining to anotherembodiment of the invention;

FIG. 6 is a flow diagram relating to the embodiment of FIG. 5;

FIG. 7 is a trace of patient flow rate versus time illustrating anotherembodiment of the invention;

FIG. 8 is a schematic illustration of a circuit according to anembodiment of the invention;

FIG. 9 is a trace of flow rate versus time showing processed andunprocessed flow rate signals;

FIGS. 10a and 10 b are illustrations of the operation of controlalgorithms according to other embodiments of the invention;

FIGS. 11a and 11 b are illustrations of the operation of still othercontrol algorithms according to other embodiments of the invention; and

FIG. 12 illustrates operation of yet another control algorithm accordingto the invention.

There is generally indicated at 10 in FIG. 1 an apparatus according toone presently preferred embodiment of the instant invention and shown inthe form of a functional block diagram. Apparatus 10 is operableaccording to a novel process which is another aspect of the instantinvention for delivering breathing gas such as air alternately atrelatively higher and lower pressures (i.e., equal to or above ambientatmospheric pressure) to a patient 12 for treatment of the conditionknown as sleep apnea.

Apparatus 10 comprises a gas flow generator 14 (e.g., a blower) whichreceives breathing gas from any suitable source, a pressurized bottle 16or the ambient atmosphere, for example. The gas flow from flow generator14 is passed via a delivery conduit 20 to a breathing appliance such asa mask 22 of any suitable known construction which is worn by patient12. The mask 22 may preferably be a nasal mask or a full face mask 22 asshown. Other breathing appliances which may be used in lieu of a maskinclude nasal cannulae, an endotracheal tube, or any other suitableappliance for interfacing between a source of breathing gas and apatient, consistent with the desired effect to be achieved through useof the apparatus 10.

The mask 22 includes a suitable exhaust port means, schematicallyindicated at 24, for exhaust of breathing gases during expiration.Exhaust port 24 preferably is a continuously open port which imposes asuitable flow resistance upon exhaust gas flow to permit a pressurecontroller 26, located in line with conduit 20 between flow generator 14and mask 22, to control the pressure of air flow within conduit 20 andthus within the airway of the patient 12. For example, exhaust port 24may be of sufficient cross-sectional flow area to sustain a continuousexhaust flow of approximately 15 liters per minute. The flow via exhaustport 24 is one component, and typically the major component of theoverall system leakage, which is an important parameter of systemoperation. In an alternative embodiment to be discussed hereinbelow, ithas been found that a non-rebreathing valve may be substituted for thecontinuously open port 24.

The pressure controller 26 is operative to control the pressure ofbreathing gas within the conduit 20 and thus within the airway of thepatient. Pressure controller 26 is located preferably, although notnecessarily, downstream of flow generator 14 and may take the form of anadjustable valve which provides a flow path which is open to the ambientatmosphere via a restricted opening, the valve being adjustable tomaintain a constant pressure drop across the opening for all flow ratesand thus a constant pressure within conduit 20.

Also interposed in line with conduit 20, preferably downstream ofpressure controller 26, is a suitable flow transducer 28 which generatesan output signal that is fed as indicated at 29 to a flow signalconditioning circuit 30 for derivation of a signal proportional to theinstantaneous flow rate of breathing gas within conduit 20 to thepatient.

It will be appreciated that flow generator 14 is not necessarily apositive displacement device. It may be, for example, a blower whichcreates a pressure head within conduit 20 and provides air flow only tothe extent required to maintain that pressure head in the presence ofpatient breathing cycles, the exhaust opening 24, and action of pressurecontroller 26 as above described. Accordingly, when the patient isexhaling, peak exhalation flow rates from the lungs may far exceed theflow capacity of exhaust port 24. As a result, exhalation gas backflowswithin conduit 20 through flow transducer 28 and toward pressurecontroller 26, and the instantaneous flow rate signal from transducer 28thus will vary widely within a range from relatively large positive(i.e. toward the patient) flow to relatively large negative (i.e. fromthe patient)-flow.

The instantaneous flow rate signal from flow signal conditioningcircuitry 30 is fed as indicated at 32 to a decision module 34, a knowncomparator circuit for example, and is additionally fed as indicated at36 to a low pass filter 38. Low pass filter 38 has a cut-off frequencylow enough to remove from the instantaneous flow rate input signal mostvariations in the signal which are due to normal breathing. Low passfilter 38 also has a long enough time constant to ensure that spurioussignals, aberrant flow patterns and peak instantaneous flow rate valueswill not dramatically affect system average flow. That is, the timeconstant of low pass filter 38 is selected to be long enough that itresponds slowly to the instantaneous flow rate signal input.Accordingly, most instantaneous flow rate input signals which could havea large impact on system average flow in the short term have a muchsmaller impact over a longer term, largely because such instantaneousflow rate signal components will tend to cancel over the longer term.For example, peak instantaneous flow rate values will tend to bealternating relatively large positive and negative flow valuescorresponding to peak inhalation and exhalation flow achieved by thepatient during normal spontaneous breathing. The output of low passfilter 38 thus is a signal which is proportional to the average flow inthe system, and this is typically a positive flow which corresponds toaverage system leakage (including flow from exhaust 24 ) since, asnoted, inhalation and exhalation flow cancel for all practical purposes.

The average flow signal output from the low pass filter 38 is fed asindicated at 40 to decision circuitry 34 where the instantaneous flowrate signal is continually compared to the system average flow signal.The output of the decision circuitry 34 is fed as a drive signalindicated at 42 to control the pressure controller 26. The pressuremagnitude of breathing gas within conduit 20 thus is coordinated withthe spontaneous breathing effort of the patient 12, as follows.

When the patient begins to inhale, the instantaneous flow rate signalgoes to a positive value above the positive average flow signal value.Detection of this increase in decision circuitry 34 is sensed as thestart of patient inhalation. The output signal from decision circuitry34 is fed to pressure controller 26 which, in response, provides higherpressure gas flow within conduit 20 and thus higher pressure within theairway of the patient 12. This is the higher magnitude pressure value ofour bi-level CPAP system and is referred to hereinbelow as IPAP(inhalation positive airway pressure). During inhalation, the flow ratewithin conduit 20 will increase to a maximum and then decrease asinhalation comes to an end.

At the start of exhalation, air flow into the patient's lungs is nil andas a result the instantaneous flow rate signal will be less than theaverage flow rate signal which, as noted is a relatively constantpositive flow value. The decision circuitry 34 senses this condition asthe start of exhalation and provides a drive signal to pressurecontroller 26 which, in response, provides gas flow within conduit 20 ata lower pressure which is the lower magnitude pressure value of thebi-level CPAP system, referred to hereinbelow as EPAP (exhalationpositive airway pressure). As has been noted hereinabove the range ofEPAP pressures may include ambient atmospheric pressure. When thepatient again begins spontaneous inhalation, the instantaneous flow ratesignal again increases over the average flow rate signal, and thedecision circuitry once again feeds a drive signal to pressurecontroller 26 to reinstitute the IPAP pressure.

System operation as above specified requires at least periodiccomparison of the input signals 32 and 40 by decision circuitry 34.Where this or other operations are described herein as continual, thescope of meaning to be ascribed includes both continuous (i.e.uninterrupted) and periodic (i.e. at discrete intervals).

As has been noted, the system 10 has a built-in controlled leakage viaexhaust port 24 thus assuring that the average flow signal normally willbe at least a small positive flow, although in some circumstances suchas when oxygen is added to the gas flow, the average flow may be anegative value. During inhalation, the flow sensed by the flowtransducer will be the sum of exhaust flow via port 24 and all othersystem leakage downstream of transducer 28, and inhalation flow withinthe airway of the patient 12. Accordingly, during inhalation theinstantaneous flow rate signal as conditioned by conditioning module 30,will reliably and consistently reflect inhalation flow exceeding theaverage flow rate signal. During exhalation, the flow within conduit 20reverses as exhalation flow from the lungs of the patient far exceedsthe flow capacity of exhaust port 24. Accordingly, exhalation airbackflows within conduit 20 past transducer 28 and toward pressurecontroller 26. Since pressure controller 26 is operable to maintain setpressure, it will act in response to flow coming from both the patientand the flow generator to open an outlet port sufficiently toaccommodate the additional flow and thereby maintain the specified setpressure as determined by action of decision circuitry 34.

In both the inhalation and exhalation cycle phases, the pressure of thegas within conduit 20 exerts a pressure within the airway of the patientto maintain an open airway and thereby alleviate airway constriction.

In practice, it may be desirable to provide a slight offset in theswitching level within decision circuitry 34 with respect to the averageflow rate signal, so that the system does not prematurely switch fromthe low pressure exhalation mode to the higher pressure inhalation mode.That is, a switching setpoint offset in the positive direction fromsystem average flow may be provided such that the system will not switchto the IPAP mode until the patient actually exerts a significantspontaneous inspiratory effort of a minimum predetermined magnitude.This will ensure that the initiation of inhalation is completelyspontaneous and not forced by an artificial increase in airway pressure.A similar switching setpoint offset lay be provided when in the IPAPmode to ensure the transition to the lower pressure EPAP mode will occurbefore the flow rate of air into the lungs of the patient reaches zero(i.e. the switch to EPAP occurs slightly before the patient ceasesinhalation.) This will ensure that the patient will encounter no undueinitial resistance to spontaneous exhalation.

From the above description, it will be seen that a novel method oftreating sleep apnea is proposed according to which the airway pressureof the patient is maintained at a higher positive pressure duringinspiration and a relatively lower pressure during expiration, allwithout interference with the spontaneous breathing of the patient. Thedescribed apparatus is operable to provide such treatment for sleepapnea patients by providing a flow of breathing gas to the patient atpositive pressure, and varying the pressure of the air flow to providealternately high and low pressure within the airway of the patientcoordinated with the patient's spontaneous inhalation and exhalation.The described system can also be used with pressure support systems suchas the proportional airway pressure system described in U.S. Pat. No.5,107,830 of Younes, the entire disclosure of which is herebyincorporated herein and made a part hereof by reference.

To provide pressure control, the flow rate of breathing gas to thepatient is detected and processed to continually provide a signal whichis proportional to the instantaneous breathing gas flow rate in thesystem. The instantaneous flow rate signal is further processed toeliminate variations attributable to normal patient respiration andother causes thus generating a signal which is proportional to theaverage or steady state system gas flow. The average flow signal iscontinually compared with the instantaneous flow signal as a means todetect the state of the patient's spontaneous breathing versus averagesystem flow. When instantaneous flow exceeds the average flow, thepatient is inhaling, and in response the pressure of gas flowing to thepatient is set at a selected positive pressure, to provide acorresponding positive pressure within the airway of the patient. Whencomparison of the instantaneous flow rate signal with the average flowsignal indicates the patient is exhaling, as for example when theinstantaneous flow signal indicates flow equal to or less than theaverage flow, the pressure of breathing gas to the patient is adjustedto a selected lower pressure to provide a corresponding lower pressurewithin the airway of the patient.

In an alternative embodiment of the invention as shown in FIGS. 2 and 3,the low pass filter 38 is replaced by an estimated leak computer whichincludes a low pass filter as well as other functional elements as shownin FIG. 3. The remainder of the system as shown in FIG. 2 is similar inmost respects to the system shown in FIG. 1. Accordingly, like elementsare identified by like numbers, and the description hereinabove of FIG.1 embodiment also applies generally to FIG. 2.

By using the operative capability of the estimated leak computer 50, asdescribed hereinbelow, it is possible to adjust the reference signalwhich is fed to decision circuitry 34 on a breath by breath basis ratherthan merely relying on long term average system flow. To distinguishthis new reference signal from average system flow it will be referredto hereinbelow as the estimated leak flow rate signal or just theestimated leak signal.

As was noted hereinabove, the average system flow rate reference signalchanges very slowly due to the long time constant of the low pass filter38. This operative feature was intentionally incorporated to avoiddisturbance of the reference signal by aberrant instantaneous flow ratesignal inputs such as erratic breathing patterns. While it was possibleto minimize the impact of such aberrations on the average flow ratereference signal, the average flow signal did nevertheless change,although by small increments and only very slowly in response todisturbances. Due to the long time constant of the low pass filter, suchchanges in the reference signal even if transitory could last for a longtime.

Additionally, even a small change in the reference signal could producea very significant effect on system triggering. For example, since theobjective is to trigger the system to the IPAP mode when inhalation flowjust begins to go positive, small changes in the reference signal couldresult in relatively large changes in the breathing effort needed totrigger the system to the IPAP mode. In some instances the change inreference signal could be so great that with normal breathing effort thepatient would be unable to trigger the system. For example, if thesystem were turned on before placement of the mask on the face of thepatient, the initial free flow of air from the unattached mask couldresult in a very large magnitude positive value for initial averagesystem flow. If such value were to exceed the maximum inspiratory flowrate achieved in spontaneous respiration by the patient, the systemwould never trigger between the IPAP and EPAP modes because the decisioncircuitry would never see an instantaneous flow rate signal greater thanthe average flow rate signal, at least not until a sufficient number ofnormal breathing cycles after application of the mask to the patient tobring the reference signal down to a value more closely commensuratewith the actual system leak in operation. As has been noted, with thelow pass filter this could take a rather long time, during which timethe patient would be breathing spontaneously against a uniform positivepressure. This would not be at all in keeping with the presentinvention.

In addition to the embodiment based on a reference signal derived fromestimated leak flow rate on a breath by breath basis which is controlledtotally by spontaneous patient breathing, two further modes of operationalso are envisioned, one being spontaneous/timed operation in which thesystem automatically triggers to the IPAP mode for just long enough toinitiate patient inspiration if the system does not sense inspiratoryeffort within a selected time after exhalation begins. To accomplishthis, a timer is provided which is reset at the beginning of eachpatient inspiration whether the inspiratory cycle was triggeredspontaneously or by the timer itself. Thus, only the start ofinspiration is initiated by the timer. The rest of the operating cyclein this mode is controlled by spontaneous patient breathing and thecircuitry of the system to be described.

A further mode of operation is based purely on timed operation of thesystem rather than on spontaneous patient breathing effort, but withtimed cycles in place of spontaneous patient breathing.

Referring to FIG. 3, the estimated leak computer 50 includes the lowpass filter 38′ as well as other circuits which are operative to makecorrections to the estimated leak flow rate signal based on on-goinganalysis of each patient breath. A further circuit is provided which isoperative to adjust the estimated leak flow rate signal quickly aftermajor changes in system flow such as when the blower has been runningprior to the time when the mask is first put on the patient, or after amajor leak in the system has either started or has been shut off.

The low pass filter 38′ also includes a data storage capability whosefunction will be described hereinbelow.

The low pass filter 38′ operates substantially as described above withreference to FIG. 1 in that it provides a long term average of systemflow which is commensurate with steady state system leakage includingthe flow capacity of the exhaust port 24. This long term average isoperative in the FIG. 3 embodiment to adjust the estimated leak flowrate reference signal only when system flow conditions are changing veryslowly.

To provide breath by breath analysis and adjustment of the referencesignal, a differential amplifier 52 receives the instantaneous flow ratesignal as indicated at 54, and the estimated leak signal output from lowpass filter 38′ as indicated at 56.

The output of differential amplifier 52 is the difference betweeninstantaneous flow rate and estimated leak flow rate, or in other wordsestimated instantaneous patient flow rate. This will be clear uponconsidering that instantaneous flow is the sum of patient flow plusactual system leakage. The estimated patient flow signal output fromdifferential amplifier 52 is provided as indicated at 58 to a flowintegrator 60 which integrates estimated patient flow breath by breathbeginning and ending with the trigger to IPAP. Accordingly, anadditional input to the flow integrator 60 is the IPAP/EPAP state signalas indicated at 62. The IPAP/EPAP state signal is the same as the drivesignal provided to pressure controller 26; that is, it is a signalindicative of the pressure state, as between IPAP and EPAP, of thesystem. The state signal thus may be used to mark the beginning and endof each breath for purposes of breath by breath integration byintegrator 60.

If the estimated leak flow rate signal from low pass filter 38′ is equalto the true system leak flow rate, and if the patient's inhaled andexhaled volumes are identical for a given breath (i.e. total positivepatient flow equals total negative patient flow for a given breath),then the integral calculated by integrator 60 will be zero and noadjustment of estimated leak flow rate will result. When the integralcalculated by integrator 60 is non-zero, the integral value in the formof an output signal from integrator 60 is provided as indicated at 64 toa sample and hold module 66. Of course, even with a zero value integral,an output signal may be provided to module 66, but the ultimate resultwill be no adjustment of the estimated leak flow rate signal.

A non-zero integral value provided to module 66 is further provided tomodule 38′ as indicated at 68 with each patient breath by operativeaction of the IPAP/EPAP state signal upon module 66 as indicated at 70.The effect of a non-zero integral value provided to module 38′ is anadjustment of the estimated leak flow rate signal proportional to theintegral value and in the direction which would reduce the integralvalue towards zero on the next breath if all other conditions remain thesame.

With this system, if the patient's net breathing cycle volume is zero,and if the system leak flow rate changes, the integrator circuit willcompensate for the change in leak flow rate by incremental adjustmentsto the estimated leak flow rate within several patient breaths.

The integrator circuit 60 also will adjust the estimated leak flow ratesignal in response to non-zero net volume in a patient breathing cycle.It is not unusual for a patient's breathing volume to be non-zero. Forexample, a patient may inhale slightly more on each breath than heexhales over several breathing cycles, and then follow with a deeper orfuller exhalation. In this case, the integrator circuit would adjust theestimated leak flow rate signal as if the actual system leak rate hadchanged; however, since the reference signal correction is only aboutone tenth as large as would be required to make the total correction inone breath, the reference signal will not change appreciably over justone or two breaths. Thus, the integrator circuit accommodates bothchanges in system leakage and normal variations in patient breathingpatterns.

An end exhalation module 74 is operative to calculate another datacomponent for use in estimating the system leak flow rate as follows.The module 74 monitors the slope of the instantaneous flow rate waveform. When the slope value is near zero during exhalation (as indicatedby the state signal input 76) the indication is that the flow rate isnot changing. If the slope of the instantaneous flow rate signal waveform remains small after more than one second into the respiratoryphase, the indication is that exhalation has ended and that the net flowrate at this point thus is the leak flow rate. However, if estimatedpatient flow rate is non-zero at the same time, one component of theinstantaneous flow rate signal must be patient flow.

When these conditions are met, the circuit adjusts the estimated leakflow rate slowly in a direction to move estimated patient flow ratetoward zero to conform to instantaneous patient flow conditions expectedat the end of exhalation. The adjustment to estimated leak flow rate isprovided as an output from module 74 to low pass filter 38′ as indicatedat 80. When this control mechanism takes effect, it disables the breathby breath volume correction capability of integrator circuit 60 for thatbreath only.

The output of module 74 is a time constant control signal which isprovided to low pass filter 38′ to temporarily shorten the time constantthereof for a sufficient period to allow the estimated leak flow rate toapproach the instantaneous flow rate signal at that specific instant. Itwill be noted that shortening the low pass filter time constantincreases the rapidity with which the low pass filter output (a systemaverage) can adjust toward the instantaneous flow rate signal input.

Another component of estimated leak flow rate control is a gross errordetector 82 which acts when the estimated patient flow rate, providedthereto as indicated at 84, is away from zero for more than about 5seconds. Such a condition may normally occur, for example, when the flowgenerator 14 is running before mask 22 is applied to the patient. Thispart of the control system is operative to stabilize operation quicklyafter major changes in the leak rate occur.

In accordance with the above description, it will be seen that low passfilter 38′ acts on the instantaneous flow rate signal to provide anoutput corresponding to average system flow, which is system leakagesince patient inspiration and expiration over time constitutes a netpositive flow of zero. With other enhancements, as described, the systemaverage flow can be viewed as an estimate of leakage flow rate.

The differential amplifier 52 processes the instantaneous flow ratesignal and the estimated leak flow rate signal to provide an estimatedpatient flow rate signal which is integrated and non-zero values of theintegral are fed back to module 38′ to adjust the estimated leak flowrate signal on a breath by breath basis. The integrator 60 is reset bythe IPAP/EPAP state signal via connection 62.

Two circuits are provided which can override the integrator circuit,including end exhalation detector 74 which provides an output to adjustthe time constant of low pass filter 38′ and which also is provided asindicated at 86 to reset integrator 60. Gross error detector 82 is alsoprovided to process estimated patient flow rate and to provide anadjustment to estimated leak flow rate under conditions as specified.The output of module 82 also is utilized as an integrator reset signalas indicated at 86. It will be noted that the integrator 60 is resetwith each breath of the patient if, during that breath, it is ultimatelyoverridden by module 74 or 82. Accordingly, the multiple resetcapabilities for integrator 60 as described are required.

In operation, the system may be utilized in a spontaneous triggeringmode, a spontaneous/timed mode or a purely timed mode of operation. Inspontaneous operation, decision circuitry 34 continuously compares theinstantaneous flow rate with estimated leak flow rate. If the system isin the EPAP state or mode, it remains there until instantaneous flowrate exceeds estimated leak flow rate by approximately 40 cc per second.When this transition occurs, decision circuitry 34 triggers the systeminto the IPAP mode for 150 milliseconds. The system will then normallyremain in the IPAP mode as the instantaneous flow rate to the patientwill continue to increase during inhalation due to spontaneous patienteffort and the assistance of the increased IPAP pressure.

After the transition to the IPAP mode in each breath, a temporary offsetis added to the estimated leak flow rate reference signal. The offset isproportional to the integral of estimated patient flow rate beginning atinitiation of the inspiratory breath so that it gradually increases withtime during inspiration at a rate proportional to the patient'sinspiratory flow rate. Accordingly, the flow rate level above estimatedleak flow needed to keep the system in the IPAP mode during inhalationdecreases with time from the beginning of inhalation and in proportionto the inspiratory flow rate. With this enhancement, the longer aninhalation cycle continues, the larger is the reference signal belowwhich instantaneous flow would have to decrease in order to trigger theEPAP mode. For example, if a patient inhales at a constant 500 cc persecond until near the end of inspiration, a transition to EPAP willoccur when his flow rate drops to about 167 cc per second after onesecond, or 333 cc per second after two seconds, or 500 cc per secondafter three seconds, and so forth. For a patient inhaling at a constant250 cc per second, the triggers would occur at 83, 167 and 250 cc persecond at one, two and three seconds into IPAP, respectively.

In this way, the EPAP trigger threshold comes up to meet the inspiratoryflow rate with the following benefits. First, it becomes easier andeasier to end the inspiration cycle with increasing time into the cycle.Second, if a leak develops which causes an increase in instantaneousflow sufficient to trigger the system into the IPAP mode, this systemwill automatically trigger back to the EPAP mode after about 3.0 secondsregardless of patient breathing effort. This would allow thevolume-based leak correction circuit (i.e. integrator 60) to act as itis activated with each transition to the IPAP mode. Thus, if a leakdevelops suddenly, there will be a tendency toward automatic triggeringrather than spontaneous operation for a few breaths, but the circuitwill not be locked into the IPAP mode.

Upon switching back to the EPAP mode, the trigger threshold will remainabove the estimated leak flow rate for approximately 500 milliseconds toallow the system to remain stable in the EPAP mode without switchingagain while the respective flow rates are changing. After 500milliseconds, the trigger threshold offset is reset to zero to await thenext inspiratory effort.

The normal state for the circuit is for it to remain in the EPAP modeuntil an inspiratory effort is made by the patient. The automaticcorrections and adjustments to the reference signal are effective tokeep the system from locking up in the IPAP mode and to preventauto-triggering while at the same time providing a high level ofsensitivity to inspiratory effort and rapid adjustment for changing leakconditions and breathing patterns.

In the spontaneous/timed mode of operation, the system performs exactlyas above described with reference to spontaneous operation, except thatit allows selection of a minimum breathing rate to be superimposed uponthe spontaneous operating mode. If the patient does not make aninspiratory effort within a predetermined time, the system willautomatically trigger to the IPAP mode for 200 milliseconds. Theincreased airway pressure for this 200 milliseconds will initiatepatient inspiration and provide sufficient time that spontaneous patientflow will exceed the reference signal so that the rest of the cycle maycontinue in the spontaneous mode as above described. The breaths perminute timer is reset by each trigger to IPAP whether the transition wastriggered by the patient or by the timer itself.

In the timed operating mode, 411 triggering between IPAP and EPAP modesis controlled by a timer with a breaths per minute control being used toselect a desired breathing rate from, for example, 3 to 30 breaths perminute. If feasible, the selected breathing rate is coordinated to thepatient's spontaneous breathing rate. The percent IPAP control is usedto set the fraction of each breathing cycle to be spent in the IPAPmode. For example, if the breaths per minute control is set to 10breaths per minute (6 seconds per breath) and the percent IPAP controlis set to 33%, then the flow generator will spend, in each breathingcycle, two seconds in IPAP and four seconds in EPAP.

FIG. 4 illustrates a control panel for controlling the system abovedescribed and including a function selector switch which includesfunction settings for the three operating modes of spontaneous,spontaneous/timed, and timed as above described. The controls forspontaneous mode operation include IPAP and EPAP pressure adjustmentcontrols 90 and 92, respectively. These are used for setting therespective IPAP and EPAP pressure levels. In the spontaneous/timed modeof operation, controls 90 and 92 are utilized as before to set IPAP andEPAP pressure levels, and breaths per minute control 94 additionally isused to set the minimum desired breathing rate in breaths per minute. Inthe timed mode of operation, controls 90, 92 and 94 are effective, andin addition the percent IPAP control 96 is used to set the timepercentage of each breath to be spent in the IPAP mode.

Lighted indicators such as LED's 96, 98 and 100 are also provided toindicate whether the system is in the IPAP or EPAP state, and toindicate whether in the spontaneous/timed mode of operation theinstantaneous state of the system is spontaneous operation or timedoperation.

An alternative embodiment of the invention contemplates detecting thebeginning of inspiration and expiration by reference to a patient flowrate wave form such as shown in FIG. 5. Comparison of instantaneous flowrate with average flow rate as set forth hereinabove provides asatisfactory method for determining whether a patient is inhaling orexhaling; however, other means for evaluating instantaneous flow ratecan also be used, and these may be used alone or in combination withaverage flow rate.

For example, FIG. 5 shows a typical patient flow rate wave form withinspiration or inhalation flow shown as positive flow above base line Band exhalation flow shown as negative flow below base line B. The flowrate wave form may thus be sampled at discrete time intervals. Thecurrent sample is compared with that taken at an earlier time. Thisapproach appears to offer the benefit of higher sensitivity to patientbreathing effort in that it exhibits less sensitivity to errors in theestimated system leakage, as discussed hereinabove.

Normally, the estimated inspiratory and expiratory flow rate wave formswill change slowly during the period beginning after a few hundredmilliseconds into the respective inspiratory and expiratory phases, upuntil the respective phase is about to end. Samples of the flow ratewave form are taken periodically, and the current sample is comparedrepeatedly with a previous sample. During inspiration, if the magnitudeof the current sample is less than some appropriate fraction of thecomparison sample, then the inspiration phase is deemed to be finished.This condition thus can be used to trigger a change to the desiredexhalation phase pressure. The same process can be used duringexhalation to provide a trigger condition for the changeover toinhalation pressure.

FIG. 6 illustrates a flow diagram for a suitable algorithm showingrepeated sampling of the patient flow rate wave form at time intervalsΔt, as shown at 110, and repeated comparison of the flow at current timet with a prior sampling, for example the flow at time t-4 as indicatedat 112. Of course, the time increment between successive samplings issmall enough that the comparison with the value observed in the fourthprior sampling event covers a suitably small portion of the flow ratewave form. For example, as shown in FIG. 5 the designations for flowsampled at time t, the sample value taken four sampling events prior attime t-4, and the time interval Δt illustrate one suitable proportionaterelationship between 66 t and the time duration of a patient breath.This is., however, only an illustration. In fact, electronic technologysuch as disclosed hereinabove would be capable of performing thisalgorithm using much smaller Δt intervals, provided other parameters ofthe algorithm, such as the identity of the comparison sample and/or theproportion relationship between the compared samples, are adjustedaccordingly.

Continuing through the flow chart of FIG. 6, if the patient's state ofbreathing is inspiration (114), the current flow sample value iscompared with the sample value observed t-4 sampling events prior. Ifthe current flow sample value is less than, for example, 75% of thecomparison flow value (116), the system triggers a change in state tothe exhalation phase (118). If the current flow sample value is not lessthan 75% of the comparison value (120), no change in breathing state istriggered.

If the breathing state is not inspiration, and the current flow samplevalue is less than 75% of the selected comparison value (122), then anexpiration phase, rather than an inspiration phase, is completed and thesystem is triggered to change the state of breathing to inspiration(124). If the specified flow condition is not met, again no change inbreathing phase is triggered (126).

The routine characterized by FIG. 6 repeats continuously as indicated at128, each time comparing a current flow sample with the fourth (or othersuitable) prior sample to determine whether the system should triggerfrom one breathing phase to the other.

Of course, it will be understood that the embodiment set forth abovewith reference to FIGS. 5 and 6 does not trigger changes in breathingper se within the context of the invention as set forth in the balanceof the above specification. Rather, in that context the FIG. 5 and 6embodiment merely detects changes between inspiration and expiration inspontaneous patient breathing, and triggers the system to supply thespecified higher or lower airway pressure as set forth hereinabove.

Common problems resulting from insufficient ventilator triggeringsensitivity include failure of the ventilator to trigger to IPAP uponexertion of inspiratory effort by the patient, and failure of theventilator in IPAP to trigger off or to EPAP at the end of patientinspiratory effort. Patient respiratory effort relates primarily to theinspiratory portion of the respiration cycle because in general only theinspiratory part of the cycle involves active effort. The expiratoryportion of the respiration cycle generally is passive. A furthercharacteristic of the respiratory cycle is that expiratory flow rategenerally may reach zero and remain at zero momentarily before the nextinspiratory phase begins. Therefore the normal stable state for apatient-triggered ventilator should be the expiratory state, which isreferred to herein as EPAP.

Although the apparatus as above described provides sensitive triggeringbetween the IPAP and EPAP modes, triggering can be still furtherimproved. The primary objectives of improvements in triggeringsensitivity generally are to decrease the tendency of a system totrigger automatically from EPAP to IPAP in the absence of inspiratoryeffort by the patient, and to increase system sensitivity to the end ofpatient inspiratory effort. These are not contradictory objectives.Decreasing auto-trigger sensitivity does not necessarily also decreasetriggering sensitivity to patient effort.

Generally, for a patient whose ventilatory drive is functioningnormally, the ideal ventilator triggering sensitivity is represented byclose synchronization between changes of state in the patient'srespiration (between inspiration and expiration), and the correspondingchanges of state by the ventilator. Many conventional pressure supportventilators which trigger on the basis of specified flow or pressurethresholds can approach such close synchronization only under limitedconditions.

In the system described hereinabove, where triggering is based onspecified levels of estimated patient flow rate, close synchronizationbetween system state changes and patient respiration state changes canbe readily achieved. This is especially true for patients having nosevere expiratory flow limitation as the flow rate reversals for suchpatients typically may correspond quite closely with changes in thepatient respiratory effort; however, flow rate reversals may notnecessarily correspond precisely to changes in patient respiratoryeffort for all patients. For example, those patients experiencingrespiratory distress or those being ventilated with pressure supportventilation may not exhibit the desired close correspondence betweenpatient effort and breathing gas supply flow rate reversals.

In addressing improvements in synchronized triggering based on flowrate, at least three different respiratory cycles are considered. Thefirst is the patient respiratory cycle as indicated by patient effort.If the patient makes any effort, it must include inspiratory effort.Patient expiration may be active or passive.

The second cycle is the ventilator respiratory cycle, that is, the cycledelivered by the ventilator. This cycle also has an inspiratory phaseand an expiratory phase, but the ventilator may or may not besynchronized with the corresponding phases of the patient respiratorycycle.

The third cycle is the flow respiratory cycle, as indicated by thedirection of flow to or from the patient's airway. Flow passes from theventilator into the patient on inspiration and out of the patient onexpiration. If one looks at this cycle only, the inspiratory/expiratorychanges could be identified by the zero flow crossings between positiveand negative flow.

Ideally, if the patient's breathing drive is competent, the flowrespiratory cycle should be synchronized with the patient respiratorycycle, with the assistance of the ventilator. One important objective ofthe triggering described here is to synchronize the ventilator cyclewith the patient cycle so that the flow cycle is coordinated withpatient inspiratory effort. With flow triggering the pressure deliveredby the ventilator affects the flow, in addition to the effect on flow ofthe patient's own efforts. This is another reason why zero flowcrossings may not necessarily be good indicators of changes in patienteffort.

As described, flow triggering provides certain advantages not availablewith some other modes of ventilator triggering. For example, inventilators which use pressure variation for triggering, when thepatient makes an inspiratory effort lung volume begins to increase. Thisvolume change causes a pressure drop, which in some conventional closedcircuit ventilators is sensed by the ventilator triggering system toinitiate breathing gas delivery. With flow triggering there is no needfor a pressure drop to occur in the breathing circuit. The triggeringalgorithm may still require the patient lung volume to increase, butsince the pressure in the ventilator circuit does not have to drop thepatient expends less energy in the inspiratory effort to achieve a givenvolume change.

Additionally, in an open circuit, that is one in which exhaled gases areflushed from the system via a fixed leak, it would be difficult togenerate the required pressure drop for pressure based triggering uponexertion of small patient inspiratory effort. That is, with an opencircuit the magnitude of patient effort required to generate even a verysmall pressure drop would be considerable. Thus, it would becorrespondingly quite difficult to achieve sufficient sensitivity in apressure based triggering system to provide reliable triggering if usedin conjunction with an open circuit. In general, the open circuit issimpler than one with a separate exhaust valve and exhaust tubing andtherefore simpler to use and less costly, and for these and otherreasons may often be preferred over a closed circuit.

In a presently preferred triggering scheme for the above described flowrate triggering system, the flow rate signal must exceed a thresholdvalue equivalent to 40 cc per second continuously for 35 milliseconds inorder to trigger the system to IPAP. If the flow rate signal drops belowthe specified threshold value during the specified time interval, thethreshold timer is reset to zero. These threshold values permittriggering with a minimal patient volume change of only 1.4 cc, whichrepresents much improved sensitivity over some prior pressure triggeredventilators.

In reality, however, such minimal patient volume change representsoverly sensitive triggering which would not be practical in actualpractice. Noise in the flow rate signal would continually reset thethreshold timer during intervals when the patient flow rate is veryclose to 40 cc per second. Therefore, it is believed a minimum volumechange for consistent triggering of the described system would be about3 cc., although this too is considered to be a very sensitive triggeringlevel. It is noted, however, that pressure signal noise can also occurin prior pressure triggered ventilators, thus creating triggeringsensitivity problems. Due to this and other details of their operationthe theoretical maximum sensitivity of such prior ventilators is betterthan the sensitivity which can actually be achieved in operation.

As one example of an operating detail which can adversely affectsensitivity in prior ventilators, if the ventilator response time isslow, the resulting time delay between the time when a triggering timerequirement is satisfied and the time when gas supply pressure actuallybegins to rise will adversely affect sensitivity. During the time delay,the patient is continuing to make inspiratory effort and thus will bedoing much greater actual work than that needed to trigger theventilator, and much greater work than would be done if the ventilatorresponded more quickly. Another example occurs during rapid breathingwhen a patient begins to make an inspiratory effort before exhalationflow rate drops to zero. In this case, an additional delay occurs untilflow reverses and the exhalation valve has time to close.

Triggering to EPAP upon exhalation also can present problems. A typicalprior pressure support ventilator will trigger off at the end ofinspiration based on the patient's flow decreasing past a flowthreshold, or on timing out of a timer, for example. Thetrigger-to-exhalation flow threshold can be either a fixed value or afraction of peak flow, for example 25% of peak flow. As one example ofthe sort of problem which can affect triggering to exhalation in priorventilators, a patient with moderately severe COPD (chronic obstructivepulmonary disease) will exhibit comparatively high respiratory systemresistance and perhaps low respiratory system elastance. This willresult in a comparatively long time constant for the patient, the timeconstant being the time taken for exhalation flow to decay to ¼ of peakflow or other suitable selected proportion of peak flow. The long timeconstant for the COPD patient means that flow will drop slowly duringinspiration.

The only way this patient will be able to trigger the ventilator toexhalation will be in increase exhalatory effort. That is, the patientmust actively decrease the flow rate by respirator muscle effort totrigger to exhalation. Further, because of the patient's large timeconstant, there may be considerable exhalation flow when the patient'sinspiratory cycle beings, and the trigger to IPAP thus will be delayed,to the detriment of ventilator sensitivity.

To achieve greater sensitivity to patient effort in an improvedembodiment of the invention, the disclosed system may be provided withapparatus to continually monitor multiple conditions rather thanmonitoring only a single condition as represented by the abovedescription pertinent to element 82. In the improved embodiment,inspiratory time is limited to no more than about 3 seconds under normalconditions, and normal exhalation flow rate is required to be nearlyzero by about 5 seconds after the beginning of exhalation.

To employ the first of these conditions the system monitors estimatedpatient flow rate and requires its value to cross zero at least onceapproximately every 5 seconds. To monitor this condition, an input isrequired from the output of a comparator which compares patient flowrate with a zero value. The comparator is a bistable device whichprovides one output (i.e. true) when patient flow is greater than zero,and an alternative output (i.e. false) when patient flow is less thanzero. If the comparator output is either true or false continuously forlonger than about 5 seconds, a signal is generated which triggers thesystem to EPAP.

The second condition is the monitoring of the IPAP/EPAP state to see ifit remains in EPAP for more than about 5 seconds. If it does, the volumewill be held at zero until there is another valid trigger to IPAP. Thisfunction prevents the volume signal from drifting away from zero whenthere is an apnea. After 5 seconds in EPAP volume should normally bezero.

Further improvements in triggering sensitivity require consideration ofmore than flow rate levels for determination of suitable triggeringthresholds. In particular, changes in the shape or slope of thepatient's flow rate wave form can be utilized to better synchronize aventilation system with spontaneous respiratory effort exerted by thepatient. Some such considerations are discussed hereinabove withreference to FIGS. 5 and 6.

A typical flow rate tracing for a normal patient experiencing relaxedbreathing is shown in FIG. 7. The simplest approach for synchronizingventilator triggering to spontaneous patient respiration has often beento use the zero flow points 129 as the reference for triggering betweenEPAP and IPAP; however, this approach does not address the potentialproblems. First, as noted above, flow rate reversals may not necessarilycorrespond to changes in patient spontaneous breathing effort for allpatients. Second, the flow signal is not necessarily uniform. Concerningthis problem in particular, small flow variations corresponding to noisein electronic signals will occur due to random variations in system andairway pressure caused by flow turbulence or random variations in thepressure control system. Additional flow rate noise results from smalloscillations in the breathing gas flow rate due to changes in bloodvolume within the patient's chest cavity with each heart beat. Third,since pressure support commonly is an all-or-nothing mode ofventilation, and since the respiratory systems of all patients willexhibit an elastance component, an inappropriate trigger to IPAP willalways result in delivery of some volume of breathing gas unless thepatient's airway is completely obstructed. Thus, reliance on zero flowor flow reversals for triggering will in many instances result in lossof synchronization between a patient inspiratory effort and ventilatordelivery of breathing gas.

Referring to FIG. 7, there is shown a flow rate versus time trace forthe respiratory flow of a typical, spontaneously breathing person. Timeadvances from left to right on the horizontal axis and flow rate variesabout a zero-value base line on the vertical axis. Of course, in thedescription below referring to FIG. 7 it is to be understood that theillustrated flow rate versus time trace is an analog for spontaneousbreathing by a patient and for any detectors or sensors, whether basedon mechanical, electronic or other apparatus, for detecting patient flowrate as a function of time.

It will be noted on reference to FIG. 7 that relatively sharp breaks orslope changes occur in the flow rate trace when patient inhalationeffort begins and ends. Thus, although the slope of the flow rate tracemay vary widely from patient to patient depending on the patientresistance and elastance, for any given patient it will tend to berelatively constant during inhalation and exhalation. The relativelysudden increase in flow trace slope when an inspiratory effort begins,and a corresponding sudden decrease in slope when inspiratory effortends, can be used to identify IPAP and EPAP trigger points. Since theabrupt changes in flow trace slope at the beginning and end ofinhalation correspond with the beginning and end of inspiratory effort,the flow trace slope or shape may be used to trigger the system betweenthe IPAP and EPAP state in reliance on changes in patient effort ratherthan on relatively fixed flow thresholds.

The basic algorithm for operation of this further improved triggeringsystem is as follows. The input to the system is the current flow rateat any point on the flow trace, for example point 130 in FIG. 7. Theflow value is differentiated to get the corresponding current slope 132of the flow function at time T. Of course, the slope corresponds to theinstantaneous rate of change of flow rate.

The slope 132 is scaled by a time factor Δt, and is then added to thecurrent flow value. This gives a prediction of the flow 134 at a futuretime t+Δt, based on the current slope 132 of the flow trace. As noted,the magnitude of time interval Δt is determined by the selected scalefactor by which slope 132 is scaled. The resulting flow prediction 134assumes, effectively, a uniform rate of change for the flow throughoutthe period Δt. It will be understood that in FIG. 7 the magnitude oftime intervals 66 is exaggerated, and is different for the end ofinspiration and beginning of inspiration changes, to facilitate clearillustration. In practice, Δt may be approximately 300 milliseconds, forexample.

The accuracy of the predicted future flow rate value 134 will dependupon the actual, varying slope of the flow trace during time interval Δtas compared to the presumed uniform slope 132. To the extent that theflow trace during interval At deviates from slope 132, the actual flowvalue 136 at the end of time interval Δt will vary from the predictedflow value 134. Accordingly, it may be seen that the effectiveness ofthis algorithm depends upon the actual deviation in the flow trace fromslope 132 during any given Δt time interval being small except when thepatient makes a significant change in respiratory effort.

To further modify the predicted flow 134, an offset factor may be addedto it to produce an offset predicted value 138. The offset is negativeduring inspiration so that the predicted flow at time t+Δt will be heldbelow the actual flow at t+Δt except when the flow trace changes in thedecreasing flow direction, such as would occur at the end of inspirationwhen patient effort ceases or even reverses.

The above description refers to an end-of-inspiration trigger to EPAP.In an entirely similar fashion, as indicated generally at 140 in FIG. 7,a trigger to IPAP can be governed by prediction of a future flow ratebased on differentiation of an instantaneous flow rate 139 plus anoffset factor to provide a predicted flow rate value 142. It is notedthat in this instance the offset factor will be a positive offset sinceit is an abruptly increasing flow rate that one would wish to detect. Inthis instance, when actual flow rate increases sufficiently during theΔt interval to exceed predicted value 142, for example as indicated at144, the system will trigger to IPAP. Thus, at any point in the flowcycle when actual flow after a predetermined time interval Δt differssufficiently from the predicted, offset flow value based on flow traceslope at time T, the system is triggered to IPAP or EPAP depending uponthe direction, either positive or negative, in which actual flow variesfrom the predicted flow.

The flow rate prediction based on current flow trace slope, and thecomparison of the predicted flow rate with actual flow rate asdescribed, is repeated at a high rate to generate a continuous or nearlycontinuous stream of actual flow-to-predicted flow comparisons. Forexample, in an analog system the process is continuous, while in adigital system the process is repeated every 10 milliseconds or faster.In the resulting locus of predicted flow value points, mostactual-to-predicted flow comparisons do not result in triggering becauseactual flow rate does not deviate sufficiently from the predicted rate.Only at the inspiratory/expiratory changes does this occur. The resultis a triggering method which, because it is based on flow predictions,does not require changes in patient inspiratory effort to achievetriggering in synchronism with spontaneous patient respiration. Ofcourse, the described differentiation technique of flow prediction isbut one example of a suitable flow wave shape triggering algorithm.

A schematic diagram of one analog circuit embodying elements of theabove-described improved trigger system is shown in FIG. 8. Input signal146 is estimated patient flow, although it may alternatively be totalflow. Use of estimated flow, in accordance with other aspects of theinvention as described hereinabove, allows for further improvedtriggering if the system includes an unknown leak component. The inputsignal 146 is differentiated by inverting differentiator 148. Thediodes, switches and operational amplifier group indicated generally at150 are included for practical considerations. They form a switchablepolarity ideal diode that is used to clip the large positive derivativeof the flow trace at the beginning of inspiration, and the largenegative derivative at the beginning of expiration. These derivativesare usually large, especially when ventilator pressure changes occur. Ifthe large derivatives are not clipped, they can interfere with triggercircuit operation in the early part of each respiratory phase.

The differentiated input signal is scaled by feedback resistor 152 andinput capacitor 154. These have a time constant equal to 300milliseconds which is a preferred delay time for the delay portion ofthe circuit. The series input resistor 156 limits high frequency noisewhich is outside the range of breathing frequencies of interest. Thedifferentiated input signal is subtracted from the flow signal in adifference amplifier 172 since the derivative circuit output is invertedand a sum is needed.

The requisite delay is produced using two fifth order, switchedcapacitor Bessel low pass filters 158. The Bessel filter has a linearphase shift with frequency which has the effect of providing a timedelay. Since the time delay depends on the order of the filter as wellas the frequency response, two filter sections are needed to provide ahigh enough cutoff frequency with the 300 millisecond time delay. The300 millisecond delay was determined experimentally. This value seemsappropriate in that the time constants for respiratory muscle activityare on the order of 50 to 100 milliseconds. Thus, a delay longer than 50to 100 milliseconds would be needed to intercept flow changes caused bychanges in respiratory muscle activity.

A 555 type timer chip 160 is set up as a 1 kHz oscillator. Along withthe RC combination 174 at the Filter input, element 160 controls thecutoff frequency of the switched capacitor filters 158. Finally, acomparator 162 with hysteresis, provided by input resistor 176 andfeedback resistors 166 and 164, changes its state based on thedifference between the input flow rate signal and the processed flowratesignal. There is a fixed hysteresis during the inspirator phase causedby resistor 166. The hysteresis characteristic of the described circuitprovides the negative offset during inspiration and the positive offsetduring expiration mentioned above.

During the expiratory phase, the hysteresis is initially greater thanthat during inspiration and is based on the magnitude of the flow signalas set by the current through resistor 164 and diode 178. The invertingamplifier 180 has a gain of 10 so that it saturates when the flow signalis greater than 0.5 volts, which corresponds to a flow of 30 liters perminute. The hysteresis is therefore almost doubled during the initialpart of the exhalation phase by virtue of the 100 k resistor 164 inparallel with the 82 k resistor 166, in contrast with the 82 k resistor166 acting alone during inspiration. That is, during exhalation theoutput of invertor 180 is positive. Therefore, diode 178 is forwardbiased and supplies current through resistor 164 in addition to thatsupplied by resistor 176.

Once the flow signal drops below the 0.5 volt level, the hysteresisdecreases linearly with flow rate to be the same as that duringinspiration. This feature prevents premature triggering to inspirationwhile gradually increasing the sensitivity of the system toward the endof exhalation. The hysteresis is somewhat dependent on the flow signalsince it is supplied through a resistor 164. With the componentsdescribed (82 k feedback resistor 166, 1.5 k input resistor 156) and a+5 volt system power supply (not shown), the hysteresis is about 90millivolts at zero flow. This corresponds with a trigger offset of 90 ccper second above the instantaneous flow signal.

FIG. 9 illustrates operation of the flow trace shape trigger circuitwith the heavy line 168 indicating the flow signal at the − input andthe lighter line 170 indicating the processed flow signal at the + inputto the system comparator 162. Line 170 shows hysteresis decreasing asflow approaches zero during exhalation.

The triggering algorithm described immediately above is referred to asshape triggering (i.e. relying on changes in the shape or slope of theflow trace) to distinguish it from the flow triggering algorithmdescribed earlier. FIGS. 10A and 10B illustrate the behavior of twoventilator triggering algorithms when used in a simulation of aspontaneously breathing patient with a large respiratory time constant.In each of FIGS. 10A and 10B, the uppermost trace represents estimatedpatient flow, the center trace represents patient respiratory effort,and the bottom trace represents ventilator pressure generated at themask which interfaces with the patient airway. The letter I indicatespatient inspiratory effort in each respiratory cycle.

As may be seen, FIG. 10A illustrates the sorts of triggering problemswhich were described hereinabove and which have been known to occur withactual patients. Patient inspiratory effort does not consistentlytrigger the ventilator, as shown by the asynchrony between patienteffort, the estimated patient flow, and pressure generated at the mask.FIG. 10B, by contrast, illustrates ventilator response to the samesimulated patient using the above-described shape triggering algorithm,but with all other settings unchanged. A comparison of FIG. 10A with 10Breveals the improved synchronization of delivered pressure withspontaneous patient effort.

It is noted that the simulated patient on which FIGS. 10A and 10B arebased exhibits significant exhalatory flow at the beginning of aninspiratory effort. From FIG. 10B, it may be clearly seen that the shapetriggering algorithm triggers appropriately at the beginning and end ofeach and every inspiratory effort even though the inspiratory effortbegins while there is still exhalatory flow. Thus, the shapes triggeringalgorithm achieves the objective of synchronizing ventilator triggeringwith changes in patient effort.

FIGS. 11A and 11B are similar to FIGS. 10A and 10B, respectively, butthe simulated patient has a shorter respiratory system time constant.FIG. 11A illustrates, from top to bottom, estimated patient flow,patient respiratory effort, and pressure at the mask for a giventriggering algorithm, and FIG. 11B illustrates the same parameters forthe above-described flow trace shape triggering algorithm. In FIG. 11Ait is clear that even with a substantial drop in flow when patienteffort stops, the system does not trigger to EPAP until flow drops belowabout 25% of peak flow. Thus, the end of the inspiratory phase is notsynchronized with the end of patient inspiratory effort. By contrast,the shape triggering algorithm of FIG. 11B produces, for the samepatient, consistently synchronized triggering at the end of patientinspiratory effort.

As may be appreciated from the above description, a ventilator can betriggered on the basis of changes in the shape of the flow signal whichcorrespond to changes in patient respiratory effort. This triggeringalgorithm solves several vexing problems which have limited the utilityof some prior triggering algorithms such as standard pressure supporttriggering. It is also to be noted that the described flow signal shapetriggering algorithm does not rely on any calculation of system leakage,although the shape triggering algorithm can operate more reliably withan estimate of system leakage than without it.

Although the described shape triggering algorithm works to trigger bothto IPAP and to EPAP, it tends to work best for triggering to EPAPbecause normally a much larger and more abrupt change in patientrespiratory effort occurs at the end of inspiration than at thebeginning.

It will be noted further that the described shape triggering algorithmmay be used in parallel with the earlier described flow triggeringalgorithm to provide a dual triggering system in which a secondarytriggering algorithm will function in the event the primary algorithmmalfunctions. This sort of dual triggering system can operate on acontinuing breath-to-breath basis such that at any desired trigger pointthe secondary triggering algorithm will trigger the ventilator if theprimary triggering algorithm does not.

Additional improvements in leak compensation techniques are alsocontemplated by this invention. As noted in the above descriptionconcerning leak compensation, the algorithm described there relies ontwo requirements as follows: (1) the patient's inhaled and exhaledvolumes over time are the same, (and indeed if the patient's rest volumejust prior to the beginning of inspiration is the same from breath tobreath, the inhaled and exhaled volumes for each individual breath willalso be the same); (2) when the patient is inhaling, total flow isgreater than leak flow and when the patient is exhaling total flow isless than leak flow.

In the above description relating to leakage, only total patient circuitflow is actually measured even though this flow is made up of twocomponents, patient flow and leakage flow. Further, in the abovedescription concerning leakage the leak is not estimated as a functionof pressure. Rather, an average leak is calculated by integrating totalflow. Since patient inhalation volume and exhalation volume areessentially the same, the average flow over a complete respiratory cycleis generally equal to the average leak per cycle.

It is to be noted further that the leak component itself can have twocomponents, namely the known or intended leak of an exhalation port andan unknown leak component resulting from one or more inadvertent leakssuch as leakage across a mask seal or at a tubing connection. Theunknown or inadvertent leak component would be quite difficult todetermine exactly as it can be a function of both pressure and time.Therefore, to the extent it is necessary to determine this leakcomponent, its value is estimated; however, significant leak managementimprovements can also be developed without separating the leak into itsintended and inadvertent components.

Regardless of whether or not the overall leak can be characterized as afunction of time or pressure, the objective is still to adjust theestimate of leakage so that its average value is the same as the averagevalue of the true leak over integral respiratory cycles. Of course, theaverage of the true leak can be obtained as the average of total flowsince the patient flow component of average total flow is zero.

If one assumes the leak is a function of pressure, such as the leak froma WHISPER SWIVEL™ connector, we then can assume the leak will likely beless at EPAP than it is at IPAP, with an average value corresponding toa pressure between EPAP and IPAP. That is, for a leak which is afunction of pressure, the leakage rate at a lower pressure is less thanthe leakage rate at a higher pressure. This characteristic of leak flowcan be utilized in algorithms other than that described hereinabove tocompensate for system leakage.

In all systems described herein, the purpose of determining leakage isto thereby determine patient flow in an open system. Of course, a closedsystem ideally has no leaks, but in practice may have inadvertent leaks;however, in an open system gas flow from the system, whether by aregulated leak at an exhaust point or through inadvertent leaks fromtubing connections and seal interfaces, will constitute a significantpart of the total system flow. Accordingly, leakage in the patientcircuit of an open system can be calculated by using the total flow asdetected by a pneumotach. This parameter, which we refer to as raw flow,includes all flow leaving or entering the gas flow source (eg., ablower) via the patient circuit. Therefore, raw flow includes bothpatient flow and system leakage. The purpose of the described leakmanagement algorithm is to separate raw flow into patient flow andleakage components.

It is not necessary to directly measure either the patient flow orleakage component of raw flow, but consequently the results of theleakage calculations are to be regarded as estimated values. One mode ofleak management is discussed hereinabove with reference to FIGS. 1 to 5.The following algorithm description relates to alternative approaches toleak management.

For a non-pressure dependent leak, one first determines when patientinhalation begins. We refer to this point in the patient's respiratorycycle as a the breath trigger and use it as a reference point to comparepatient inspiratory and expiratory volumes. By using the beginning ofinspiration as a breath trigger point, the leakage calculation is donewith the patient's beginning lung volume at approximately the same levelfor every calculation.

The breath trigger reference point is identified by satisfaction of twoconditions as follows: (1) raw flow greater than average leak rate; and(2) patient is ready to inhale. Concerning the first of theseconditions, the comparison of raw flow with average leak rate revealswhether the patient is inhaling or exhaling because, as noted above, thenature of the raw flow component is such that raw flow exceeds systemleakage during patient inspiration and is less than system leakageduring patient expiration.

The second condition, that of determining whether the patient its readyto inhale, is assessed by reference to either of two additionalconditions as follows: (1) the patient has exhaled a predeterminedfraction of the inhaled volume, for example ¼ of inhalation volume; or(2) the patient has exhaled for more than a predetermined time, forexample 300 milliseconds. An alternative statement of the first of theseconditions is that the patient's exhaled volume is at least asignificant fraction of inspiratory volume. As to the second condition,an exhalation time of more than 300 milliseconds can be detected bycomparison of raw flow with estimated average leakage rate since, asnoted above, raw flow is less than estimated average leak rate duringexhalation. If raw flow thus is less than the average leak rate forlonger than 300 milliseconds, the condition is met. If either of theconditions 1 or 2 immediately above is met the patient is ready toinhale. Therefore, if either of the conditions is met and in additionthe raw flow is greater than average leak rate, the conditions for abreath trigger are satisfied and the described conditions thus can beused to trigger the ventilator for an inspiratory cycle.

The seeming contradiction of requiring raw flow greater than averageleak rate to satisfy one breath trigger condition, and raw flow lessthan average leak rate as one parameter that can satisfy the secondbreath trigger condition is explained as follows. The condition that rawflow is greater than average leak rate merely indicates that the patienthas begun an inhalation cycle; that is, the patient has exerted theinitial inspiratory effort that is revealed as raw flow exceedingaverage system leakage. However, as discussed hereinabove, since the rawflow signal can and does cross the average leak rate signal due tospurious noise in the raw flow signal, not all incidents of raw flowbeing greater than average leakage will denote the beginning of patientinspiratory effort. In fact, noise in the raw flow signal may typicallyresult in multiple crossings between the raw flow signal and the averageleak signal. Especially troublesome in this regard is noise causing amomentary − to + transition while the actual major condition is the flowcrossing zero due to a true inspiratory to expiratory transition.

In order to avoid multiple triggers at these points of signal crossingsthat do not denote initial patient inspiratory effort, the additionalready-to-inhale condition is imposed. Under this condition, the raw flowgreater than average leak condition is validated only if one of the twoconditions indicating the patient's respiration was just previously inan exhalatory state is satisfied. In the case of the second of theseconditions, the 300 millisecond duration of raw flow less than averageleakage corresponds to patient exhalation. Thus, the time element playsa role in the described breath trigger algorithm and must be understood.Raw flow less than average leakage indicates exhalation which thenprovides a ready-to-inhale output. A subsequent reversal of raw flow toa value greater than average leakage indicates the detection of patientinspiratory effort. The breath trigger thus consistently initiates aninspiratory cycle in synchronism with spontaneous patient respiration.

When the breath trigger occurs, the ready-to-inhale parameter is resetto await satisfaction of one of the two conditions specified above thatprovides a ready-to-inhale validation. The breath triggering process isthen repeated upon occurrence of the next patient inspiratory effortwhen raw flow again exceeds average system leakage.

The time interval between breath triggers from one breath to the nextcan be captured and the raw flow signal integrated over that timeinterval to find the raw volume or total volume for each breath. The rawvolume may then be divided by the time between breath triggers, forexample the time interval between the next successive pair of breathtriggers, to determine a recent time rate of leakage.

It is noted that the patient's respiration from one breath trigger tothe next, as represented by the integration of the raw flow signal,rises from zero volume to a maximum value comprised of inspiratoryvolume and leakage volume during inspiration. As the respiratory cyclecontinues through expiration, the continued integration of negative flowreduces the raw volume parameter progressively as the exhaled volume issubtracted. Thus, at the next breath trigger, the raw volume parameterfrom the prior breath is equal to only the leakage volume and any changein patient resting volume; however, this latter value generally isassumed to be zero.

Thus, as the raw volume parameter is really only leak volume, dividingthat value by the time duration of the breath provides an estimate ofthe leak rate for the most recent breath. By averaging the recent leakrate over time an average leak rate can be determined to provide a morestable signal. Among other benefits, this reduces any effects ofbreath-to-breath volume variation resulting from breath-to-breathchanges in patient resting volume. The number of respiratory cycles overwhich the recent leak may be averaged to determine average leak rate maybe anywhere between one and infinity, depending upon the desired balancebetween signal stability and rapid error correction. Longer termaveraging results in greater stability but slower error correction.Shorter term averaging provides less stability but quicker errorcorrection.

Finally, it is to be noted that the average leak rate whose calculationis described here was used initially in this algorithm in one of theconditions for initiating a breath trigger. The parameters used in theconditions for initiating a breath trigger thus can be continuallyupdated, used to satisfy the breath trigger conditions, and then thesuccessive breath triggers are used in turn to again update theseparameters.

A related approach to leakage analysis involves the use of a leakcomponent which is, or is assumed to be, a function of patient circuitpressure. This sort of leak analysis can be useful in such ventilationregimens as, for example, proportional assist ventilation such asdescribed in U.S. Pat. No. 5,107,830. It is important in proportionalassist ventilation to know both the average leak rate and theinstantaneous leak for the system. However, proportional assistventilation involves variation of pressure according to the level ofventilation assistance required by the patient from moment to moment.Since at least some components of system leakage can be a function ofpressure, those components will also vary more or less in synchronismwith varying patient ventilation needs.

Thus, an alternative algorithm for leak analysis that can be used in aproportional assist ventilation system and in other systems as deemedsuitable would be based on the requirement that patient flow equals rawflow, as defined hereinabove, less any known leak component and anypressure dependent leak component. Known leakage may be any leak,intentional or otherwise, having known flow characteristics, for examplethe leakage through a WHISPER SWIVEL™ or through a plateau exhalationvalve. That is, the known leak is not necessarily a fixed leak, but canalso be a function of pressure. Although such leaks of knowncharacteristics could be lumped together with leaks having unknowncharacteristics but assumed to be functions of pressure, the followinganalysis can be carried out with leaks of known characteristics includedor excluded from the leakage calculation described.

The pressure dependent leak, whether constituted of total leakage oronly a component of total leakage, is calculated as a function of systempressure from analysis of the raw flow signal. One function by which thepressure dependent leak component is related to pressure is that for anyorifice defining a pressure drop between a pressurized system and alower pressure surrounding environment, the flow through the orificewill be proportional to the square root of the pressure differenceacross the orifice multiplied by a constant K, which characterizes themechanical features of the orifice itself, that is its size, surfacesmoothness, and so forth.

To find the pressure dependent leak component, it is necessary only tocharacterize one hypothetical orifice as representing the source of allpressure dependent leakage. To find the hypothetical orifice, the breathtrigger defined above is utilized to mark the beginning and end ofpatient respiratory cycles, and the average patient circuit pressure ismeasured for each breath. The average leak rate, as calculatedhereinabove, is multiplied by the time duration of the previous breathto determine pressure dependent leak volume on a per-breath basis. Ifthe pressure dependent leak component is separated from the known leakcomponent, then the known leak component is also separated from averageleak rate before the average leak rate is used to determine pressuredependent leak volume.

The pressure dependent leak volume is then divided by the average of thesquare root of pressure for the prior breath to give the correspondingorifice characteristic K. Over the next breath, the pressure dependentleak is found by multiplying the calculated orifice characteristic K bythe square root of instantaneous pressure. This gives the pressuredependent leak rate as a function of pressure throughout the breath.This is an important parameter in such ventilation regimens as the abovecharacterized proportional assist ventilation. More generally, however,it is important for the therapist to have reliable information on systemleakage. The pressure dependent leak rate, therefore, can be useful forcharacterizing leakage in a variety of ventilation systems.

For either of the last described leak calculations, recovery routinesare desired to deal with those instances when a breath trigger does notoccur when it should due to failure of the raw flow signal to cross theaverage leak rate signal as the patient breaths. When the system leakchanges, the raw flow signal will increase if the leak increases ordecrease if the leak decreases. It is possible for this increase ordecrease to be large enough that the raw flow signal never crosses theaverage leak rate, and when this occurs there will be no breath triggerand thus no new or updated leak calculations. To recover from such anincident, an algorithm is needed to begin the leakage calculations againand thereby bring the calculated average leak rate back into line withactual system functioning. Initiation of the recovery algorithm isdetermined to be necessary if any one of four known physiological eventsoccurs as follows: (1) exhalation continues for more than 5 seconds; (2)inhalation continues for more than 5 seconds; (3) inspiratory tidalvolume is greater than 5 liters; or (4) expiratory tidal volume is lessthan −1 liter.

The ventilator system should include elements for monitoring operationto detect occurrence of any one of these events. When one such event isdetected, rediscovery of the leak is initiated by changing the averageleak rate to the raw flow signal by gradually adding a percentage of thedifference between raw flow and average leak to the average leak rate.Preferably, the percentage of difference added should be selected toprovide a time constant of approximately one second.

As the gradual increase in average leak rate in accordance with therecovery algorithm causes the average leak rate value to approach rawflow value, the increasing average leak rate ultimately will cause abreath trigger in accordance with one of the triggering algorithms asdescribed above, and the repetitive leak calculations will then resumewith each respiratory cycle.

From the above leak management algorithm it may be seen that thealgorithm may also be based on comparison of estimated patient flow to azero value, and calculate only an unknown leak as a function ofpressure. In a manner similar to the leak management algorithm abovecharacterized, a breath indicator or trigger is employed as a marker forthe start of patient inspiration. The breath trigger occurs when twoconditions are met, namely: (1) estimated patient flow is greater than0; and (2) patient is ready to inhale. The ready-to-inhale condition issatisfied when: (1) the patient has exhaled a specified portion (eg.25%) of inspiratory volume; (2) estimated patient flow is less than 0;and (3) inspiratory volume is greater than a constant represented by 0or a small positive value.

As with the above-described algorithms, the conditions requiringinspiratory volume to be both less than a percentage of expiratoryvolume and greater than a small positive value are not contradictory asthese measurements are taken at different times. First, theready-to-inhale condition must be satisfied by the specified parameters,including detection of an estimated patient flow less than zero sincethe last inhalation. Then, a breath trigger can occur when estimatedpatient flow goes positive since estimated patient flow greater than 0indicates initial patient inhalatory effort.

From one breath trigger to the next, raw flow is integrated and thesquare root of pressure is also integrated. Then at the next breathtrigger the unknown orifice is calculated by dividing the value obtainedfrom raw flow integration by the value obtained from integration of thesquare root of pressure to provide the characteristic K for the unknownorifice. As noted above, the unknown orifice is a hypothetical leaksource intended to describe the characteristic of leakage from allsources of pressure dependent leak.

Once the unknown orifice characteristic has been calculated, it ismultiplied by the square root of patient circuit pressure at any timeduring the following breath to provide a measure of the instantaneousunknown leak. Estimated patient flow is then found by subtracting theinstantaneous unknown leak from the raw flow.

In both of the algorithms which treat the system leak as the gas flowingthrough a hypothetical orifice of characteristic K, the constant K iscalculated in essentially the same way. Although the equation for thecalculation is derived differently for the two approaches, the result isthe same, as follows. $\begin{matrix}{K = \frac{\int_{0}^{T}{\left( {{total}\quad {flow}} \right){t}}}{\int_{0}^{T}{\left( \sqrt{Pressure} \right){t}}}} & {{eq}.\quad 1}\end{matrix}$

We have noted above that as one aspect of the invention, we calculateleakage as a function of gas pressure. Accordingly, the equation forcalculating the constant K can be more generally stated as follows:$\begin{matrix}{K = \frac{\int_{0}^{T}{\left( {{total}\quad {flow}} \right){t}}}{\int_{0}^{T}{{f\left( {P(t)} \right)}{t}}}} & {{eq}.\quad 2}\end{matrix}$

In cases where the system also includes a known leak component such as aleak resulting from use of a plateau valve or a WHISPER SWIVEL™ asdescribed above, the calculations can be modified to provide greateraccuracy by subtracting the known leak component from raw flow prior tointegration of raw flow to determine the unknown orifice, and also priorto utilizing raw flow to determine estimated patient flow. In this case,the integral of raw flow minus the known leak, divided by the integralof the square root of patient circuit pressure will provide theparameter K characterizing the unknown orifice attributable to only theunknown leak. That orifice characteristic K can then be multiplied bythe square root of patient circuit pressure during the subsequent breathto provide a measure of the unknown leak component. Estimated patientflow at any time then would be found by subtracting both the known leakand the calculated unknown leak from the raw flow signal.

The modified algorithm preferably also includes a recovery routine foressentially the same reasons as earlier stated. A breath trigger forthis modified algorithm depends upon estimated patient flow crossing thezero value as the patient breaths. When the leak value changes, both rawflow and estimated patient flow will increase if the leak is increasing,or decrease if the leak is decreasing. It is possible for such anincrease or decrease to be large enough that estimated patient flow willnever cross the 0 value. When this occurs no breath trigger occurs andthere will be no new or updated leak calculations unless a recoveryroutine is performed.

To correct this condition, a recovery algorithm may be utilized todetect when leakage calculation is out of control according to twolimiting parameters as follows: (1) inspiratory tidal volume is greaterthan a physiological value such as 4.5 liters; or (2) inspiratory tidalvolume is less than a non-physiological value such as −1 liter. Wheneither of these non-physiological events occurs, the leakage calculationis out of control and the system leak must be rediscovered byartificially inducing a breath trigger. Thus, the system would requireflow detecting apparatus to sense the specified limiting values ofinspiratory tidal volume and provide a breath trigger reference point.This restarts system operation essentially in the same manner asdescribed above with reference to other algorithms, and thus initiates anew leak calculation. The data used for these continued calculations caninclude data from after the change in leak magnitude so that theestimate of the unknown leak is improved. Even if the new unknown leakvalue is incorrect, continuing repeated calculations also will be basedon data from after the change in leak magnitude. Thus, the calculatedestimate of unknown leak will continue to quickly approach reliablevalues.

FIG. 12 illustrates the described leak management algorithm in action,with intentional introduction of a very large leak. In FIG. 12, the fourtraces represent, from top to bottom, pressure, estimated patient flow,estimated inspiratory volume, and estimated leak (combined known andunknown leakage). The X axis is progressing time, moving from left toright.

As can be seen from FIG. 12, for the first three respiratory cycles thealgorithm is able to determine leakage as a function of pressure. At thestart of each inspiration an unknown orifice characteristic K iscalculated. Points 1 and 2 on the estimated leak trace show thetransitions between one unknown orifice characteristic K and the next,which is an updated unknown orifice characteristic. Point 2 on theestimated leak trace shows a much smoother transition between old andnew unknown orifice characteristics than does point 1, thus indicatingthat point 2 is a transition between unknown orifice characteristics ofessentially the same value.

To test algorithm response to leakage, a very large leak was introducedas indicated at point O on the estimated patient flow trace.Predictably, the estimated patient flow and estimated inspiratory volumerise off scale very quickly. At time T, estimated inspiratory volume hasreached 4.5 liters and a breath trigger thus is forced. The resultingcalculation estimates the leak to be a very high value which, in theestimated leak portion of FIG. 12 is off scale. The resulting estimatedpatient flow is approximately 50 liters per minute where in fact itshould be zero. Tidal volume continues to rise rapidly as indicated atpoint 3 and would ultimately force another breath trigger when itreached the 4.5 liter limit. The large system leak is removed at timeT′, however, so the rise in tidal volume ceases.

With the leak removed from the system, estimated patient flow andestimated inspiratory volume rapidly drop off scale. Thus, wheninspiratory tidal volume reaches the recovery limit of −1 liter, anotherbreath trigger is forced thus leading to a new leak calculation and aresulting estimated leak lower than the previous leak calculation. Thisbrings estimated patient flow back onto scale. Although the new leakcalculation is better than the previous one, it is not by any meansperfect. With estimated patient flow based on this new unknown leakorifice characteristic, tidal volume may once again drop below the −1liter limit such as indicated at point 4 thus forcing another breathtrigger. Once again a new unknown orifice calculation is performed andused for an unknown leak calculation which results in a very good value,thus returning the system to normal ranges of operation. The system thuswas reliable in recovery from even very large changes in leak magnitude.The much smaller changes in leak magnitude that are more commonlyencountered are readily handled under this algorithm by either thenormal calculation or the recovery calculation.

Of course, we have contemplated various alternative and modifiedembodiments of the invention of which the above described are exemplaryas the presently contemplated best modes for carrying out the invention.Such alternative embodiments would also surely occur to others skilledin the art, once apprised of our invention. Accordingly, it is intendedthat the invention be construed broadly and limited only by the scope ofthe claims appended hereto.

We claim:
 1. A method of detecting changes in a respiratory state of apatient between inspiration and expiration, comprising: monitoring arespiratory gas flow of such a patient; determining, from the monitoringstep, a respiratory gas flow characteristic at selected time intervals;comparing selected pairs of said respiratory gas flow characteristics,wherein each selected pair of respiratory gas flow characteristicsincludes a predicted value of such a patient's respiratory gas flow rateat a selected time, and such a patient's actual respiratory gas flowrate at the selected time; and identifying, from the comparing step,changes in the patient's respiratory state between inspiration andexpiration.
 2. The method of claim 1, further comprising: taking aderivative of a known patient respiratory gas flow rate; and processingthe derivative to determine the predicted value of such a patient'srespiratory gas flow rate.
 3. The method of claim 2, wherein theprocessing step includes scaling the derivative with a scaling factor toobtain a scaled derivative.
 4. The method of claim 3, wherein theprocessing step includes modifying the scaled derivative with an offsetfactor to determine the predicted value of such a patient's respiratorygas flow rate.
 5. The method of claim 3, wherein the scaling factorincludes a time factor.
 6. The method of claim 5, the time factor isdetermined, at least in part, by a time span between the selected timeand a time prior to the selected time.
 7. The method of claim 1, whereinthe predicted value of such a patient's respiratory gas flow rate isdetermined from an actual patient respiratory flow rate that occurred ata time prior to the selected time.
 8. A method of determining a leakcomponent of gas flow in a respiratory gas supply system for supplyingrespiratory gas to a patient, comprising: determining a beginning and anend of an integral number of patient breaths; integrating a patientrespiratory gas flow rate throughout at least one breath to determine atotal respiratory gas volume for the at least one breath; and dividingthe total respiratory gas volume for the at least one breath by a timeduration of the at least one breath to determine a leakage rate for theat least one patient breath.
 9. The method of claim 8, furthercomprising averaging the leakage rate for the at least one patientbreath to determine an average respiratory gas leakage from such arespiratory gas supply system.